The field of the invention is nuclear magnetic resonance imaging (MRI) methods and systems. More particularly, the invention relates to the rapid acquisition of three-dimensional MR images using a shells sampling trajectory.
Any nucleus that possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant γ of the nucleus). Nuclei which exhibit this phenomena are referred to herein as “spins”.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment MZ is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the pulsed excitation signal B1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance (“NMR”) phenomena is exploited.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (Gx, Gy, and Gz) which have the same direction as the polarizing field B0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified. The acquisition of the NMR signals samples is referred to as sampling k-space, and a scan is completed when enough NMR cycles are performed to fully sample k-space.
In conventional, fully-sampled MRI, the number of acquired k-space data points is determined by the spatial resolution requirements, and the Nyquist criterion for the alias-free field of view (FOV). Images can be reconstructed, however, using a reduced number of k-space samples, or “undersampling”. The term undersampling here indicates the Nyquist criterion is not satisfied, at least in some regions of k-space. Undersampling is used for several reasons, including reduction of acquisition time, reduction of motion artifacts, achieving higher spatial or temporal resolution, and reducing the tradeoff between spatial resolution and temporal resolution. Aliasing artifacts that result from undersampling are not as severe if the violation of the Nyquist criterion is restricted to the outer part of k-space.
The time required to fully sample 3D Cartesian k-space is relatively long. Alternative non-Cartesian trajectories can provide faster coverage of k-space, and more efficient use of the gradients. When a very fast volume acquisition is required, undersampling strategies can be used in conjunction with these non-Cartesian trajectories to further reduce the scan time. The method of Lee J H, Hargreaves B A, Hu B S, Nishimura D G; Fast 3D Imaging Using Variable-Density Spiral Trajectories With Applications To Limb Perfusion, Magn. Reson. Med. 2003; 50(6): 1276-1285, uses a variable-density stack of spiral trajectories that varies the sampling density in both the kx-ky plane and the kz direction. That method preserves reasonable image quality, while reducing the acquisition time by approximately half compared to a fully-sampled acquisition. Vastly undersampled 3D projection acquisition as described by Barger V A, Block W F, Toropov Y, Grist T M, Mistretta C A, Time-Resolved Contrast-Enhanced Imaging With Isotropic Resolution and Broad Coverage Using An Undersampled 3D Projection Trajectory, Magn. Reson. Med. 2002; 48(2):297-305, has been used to increase temporal resolution and provide better dynamic information for 3D contrast-enhanced MRA. The aliasing caused by undersampling in this method often can be tolerated in angiographic applications. This is because the vessel-tissue contrast is high and the artifacts are distributed, or spread out in the image.
One such method is disclosed in U.S. Pat. No. 5,532,595 which is incorporated herein by reference. This so-called “shells” k-space sampling trajectory samples a spiral pattern in k-space around a spherical surface. A complete image acquisition is comprised of a series of such spiral sampling patterns over a corresponding series of spheres of increasing diameter. The shells k-space sampling trajectory acquires 3D data on concentric spherical surfaces in k-space. It was originally proposed in the mid-1990's, but its feasibility for image acquisition has only recently been demonstrated. The shells k-space trajectory has favorable properties for motion correction and accelerated acquisition with undersampling.
A limitation on the use of the shells sampling pattern is its sensitivity to off-resonance effects. Protons in lipids (fat) resonate at a Larmor frequency that is approximately 145 Hz/T lower than protons in water. Consequently at 1.5 T, the protons from fat are off-resonance by approximately 217 Hz. For standard Cartesian acquisitions, this causes a simple spatial shift in the frequency encoded direction, but for non-Cartesian acquisitions like shells it causes blurring. Fat suppression methods like CHESS can be used, but they add approximately 10 ms or more to the repetition time TR. For applications like contrast enhance MR angiography, a short TR time is required and CHESS fat sat is not practical. Because the T1 of fat is longer than the T1 of contrast enhanced arterial blood (e.g., approximately 250 ms vs. 50 ms at a field strength of 1.5 T), using the shortest possible TR (e.g., 5 ms or less) provides a useful degree of fat suppression due to well-known saturation effects that can be readily calculated, for example, assuming spoiled gradient echo contrast.
Reducing the readout duration during which data are acquired also reduces blurring from off-resonance effects such as the chemical shift of lipids mentioned above, and also from susceptibility variation. A standard method to reduce the readout duration is to increase the readout bandwidth (BW), but this has limitations in terms of signal-to-noise ratio (proportional to 1/√{square root over (BW)}), maximum gradient amplitude and slew rate available on any given set of gradient hardware, and peripheral nerve stimulation from increased gradient slewing.